U.S. patent number 6,224,554 [Application Number 09/282,514] was granted by the patent office on 2001-05-01 for method to measure ambient fluid pressure.
This patent grant is currently assigned to Point Biomedical Corporation. Invention is credited to Stanley R. Conston, E. Glenn Tickner.
United States Patent |
6,224,554 |
Tickner , et al. |
May 1, 2001 |
Method to measure ambient fluid pressure
Abstract
A method is provided for measuring real time ambient pressure at
a region of interest in a fluid-filled body cavity by introducing
into the cavity a composition of gas-containing microbubbles having
a predetermined fragility threshold correlating to the rupture
response of their capsules to the ambient fluid pressure and/or
applied acoustic pressure. An ultrasonic signal is applied at the
region of interest at a power level sufficient to destroy the
microbubble population having a fragility threshold below the
applied power level. The ultrasound backscatter response is
detected from the population of intact and disintegrating
microbubbles remaining at the region of interest and this
backscatter signal is correlated to predetermined acoustic response
properties to determine the ambient pressure at the region of
interest.
Inventors: |
Tickner; E. Glenn (Los Gatos,
CA), Conston; Stanley R. (San Carlos, CA) |
Assignee: |
Point Biomedical Corporation
(San Carlos, CA)
|
Family
ID: |
23081847 |
Appl.
No.: |
09/282,514 |
Filed: |
March 31, 1999 |
Current U.S.
Class: |
600/438;
600/458 |
Current CPC
Class: |
A61B
8/04 (20130101); A61B 8/481 (20130101) |
Current International
Class: |
A61B
8/00 (20060101); A61B 8/04 (20060101); A61B
008/00 () |
Field of
Search: |
;600/437,438,458
;73/1.85 |
References Cited
[Referenced By]
U.S. Patent Documents
Foreign Patent Documents
Other References
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Ventricular Performance and Systemic Hemodynamics by Study of
Aortic Root Pressure and Flow Estimates in Healthy Men, and Men
with Acute and Healed Myocardial Infarction", The Amer. J. Of
Cardiology, vol. 72. Aug. 1, 1993, 260-267. .
Bouakaz et al., "On the effect of lung filtering and cardiac
pressure on the standard properties of ultrasound contrast agent",
Ultrasonics, 36, 703-708, 1998. .
Currie et al., "Continuous Wave Deppler Dermination of Right
Ventricular Pressure: A Simultaneous Doppler--Catheterization Study
in 127 Patients," JACC vol. 6, No. 4, Oct/1985, 750-6. .
Gersh et al., "Physical criteria for measurement of left
ventricular pressure and its first derivative'", cardiovascular
Res. 1971, 5, 32-40. .
Hasegawa et al., "Acoustic radiation pressure acting on spherical
and cylindrical shells", J. Acoust. Soc. Am. 93, 1, 1/93, 154-161.
.
Himelman M.D., et al., "Noninvasive Evaluation of Pulmonary Artery
Pressure During Exercise by Saline-Enhanced Doppler
Echocardiography in Chronic Pulmonary Disease", Circulation vol.
79, No.4, Apr. 1989, 863-869. .
Ishihara et al., "New Approach to Noninvasive Manometry Based on
Pressure Dependent Resonant Shift of Elastic Microcapsules in
Ultrasonic Frequency Characteristics", Pros. Of 8th symposium on
Ultras. Elect. Tokyo 1987, Jap. J. Of Applied Physics, vol. 27,
1988, pp. 125-127. .
Kyriakides et al., "Noninvasive determination of the left
ventricular end-systolic pressure", Am.J.of Card.33, 1991, 267-274.
.
Laaban, M.D., et al., "Noninvasive Estiamtion of Systolic Pulmonary
Artery Pressure Using Doppler Echocardiography in Patients with
Chronic Obstructive Pulmonary Disease", Chest, 96, Dec. 1989,
1258-1262. .
P.A.Lewin, "Acoustic pressure amplitude thresholds for rectified
diffusion in gaseous microbubbles in biological tissue",
J.Acoust.Soc.am. 69, 3, Mar. 1981, 846-852. .
Neumann et al., "Accurate Noninvasive Estimation of Left
Ventricular End-Diastolic Pressure: Comparison with
Catheterization", J. of Amer. Soc. Of Echocardiography, vol. 11,
No.2, 1998, 126-131. .
W.L.Nyborg, "Radiation Pressure on a Small Rigid Sphere", The J. of
the Acoustic. Soc. Of Amer., May 1967, 947-952. .
Oesterle M.D., et al., "A New Method for Assessing Right-Sided
Heart Pressure Using Encapsulated Microbubbles--A Preliminary
Report", The Western J. Of Med., Oct. 1985, 463-468. .
Serruys M.D. et al., "Intracoronary Pressure and Flow Velocity with
Sensor-Tip Guidewires: A New Methodologic Approach for Assessment
of Coronary Hemodynamics Before and After Coronary Interventions",
The Amer. J.of Card.vol.71, May 20, 1993, 41D-53D. .
Vuille M.D., et al., "Effect of Static Pressure on the
Disappearance Rate of Specific Echocardiographic Contrast Agents",
J.of the Amer.Soc.of Echocardiography Jul. 8,1994, 347-354. .
Yosioka et al., "Acoustic Radiation Pressure on a Compressible
Sphere", ACUSTICA vol. 5, 1955, 167-173. .
Yosioka et al., "Acoustic Radiation Pressure on Bubbles and Their
Logarithmic Decrement", ACUSTICA, vol.5, 1955, 173-178..
|
Primary Examiner: Jaworski; Francis J.
Attorney, Agent or Firm: Fish & Richardson P.C.
Claims
What is claimed is:
1. A method for measuring ambient pressure at a region of interest
in a fluid-filled body cavity or vessel comprising the steps
of:
(a) introducing into said cavity or vessel a composition of
gas-containing microbubbles, said microbubbles having a
predetermined fragility threshold, said fragility threshold
correlating the disintegration response of said microbubbles to a
combination of fluid and applied acoustic pressure, said acoustic
pressure being applied from an ultrasonic energy producing source
and said composition having predetermined acoustic response
properties correlating to ambient pressure of a surrounding
fluid;
(b) applying an ultrasonic signal at said region of interest within
said cavity or vessel at a power level sufficient to cause acoustic
pressure to disintegrate a microbubble population, said power level
being above or equal to said predetermined fragility threshold;
(c) detecting the returned acoustic signals backscattered from the
population of disintegrating and intact microbubbles remaining at
said region of interest;
(d) correlating said returned acoustic signals to said
predetermined acoustic response properties of said composition to
determine said ambient pressure at said region of interest;
(e) repeating steps (a)-(b).
2. A method according to claim 1 further comprising the step of
displaying said correlation on a monitor of an ultrasound
scanner.
3. A method according to claim 1 wherein said microbubbles are
characterized by having a shell comprising one or more layers.
4. A method according to claim 3 wherein in step (a) the
predetermined fragility of each of said microbubbles is controlled
by selecting the thickness of one or more layers of the shell.
5. A method according to claim 3 wherein in step (a) the
predetermined fragility threshold of said microbubbles is
controlled through the use of materials in the shell of differing
moduli of elasticity.
6. A method according to claim 3 wherein in step (a) the
predetermined fragility threshold of said microbubbles is
controlled through use of materials in the shell of differing
molecular weight.
7. A method according to claim 3 wherein in step (a) the
predetermined fragility threshold of said microbubbles is
controlled through use of materials in the shell of differing hoop
strength.
8. A method according to claim 3 wherein in step (a) the
disintegration response of said microbubbles approximates a step
function at one or more pressure levels.
9. A method according to claim 1 wherein said gas in said
microbubbles comprises physiologically acceptable gas.
10. A method according to claim 9 wherein said gas comprises
nitrogen.
11. A method according to claim 9 wherein said gas comprises
air.
12. A method according to claim 9 wherein said gas comprises a
fluorocarbon compound.
13. A method according to claim 1 wherein said composition is
introduced into said cavity as a bolus.
14. A method according to claim 1 wherein said composition is
introduced to said cavity as a controlled rate infusion.
15. A method according to claim 1 wherein in step (a) the
disintegration response is a linear relationship with ambient
pressure in combination with constant ultrasound pressure.
16. A method according to claim 1 wherein said step of applying an
ultrasonic signal and said step of detection are accomplished with
an ultrasound imaging system.
17. A method according to claim 16 wherein the returned acoustic
signals of said microbubbles is detected utilizing B-mode harmonic
imaging methods.
18. A method according to claim 16 wherein the returned acoustic
signals of said microbubbles is detected utilizing power Doppler
decorrelation imaging methods.
19. A method according to claim 18 wherein the ambient fluid
pressure is determined by comparing the intensity of the returned
power Doppler decorrelation signal and the predetermined acoustic
response properties of said microbubbles.
20. A method according to claim 19 wherein the ambient fluid
pressure result is displayed in real time.
21. A method according to claim 19 wherein the ambient fluid
pressure result is displayed as a greyscale map of ambient pressure
within said body cavity. map of ambient pressure.
22. A method according to claim 19 wherein the ambient fluid
pressure result is displayed as a color map of ambient pressure
within said body cavity.
23. A method according to claim 22 wherein the colors of the color
map are coded to the amplitude of the ambient pressure.
24. A method according to claim 19 wherein the results are
displayed as a map of ambient pressure and correlated to a
predetermined electrocardiogram of the subject to display pressure
as a function of time point within the cardiac cycle.
25. A method according to claim 1 wherein said composition
comprises two or more population of microbubbles of differing
fragility thresholds.
26. A method according to claim 1 wherein the response of said
microbubbles is detected by way of interrogation with an ultrasound
signal of a first low power signal to establish a baseline and a
second higher power signal to disintegrate the microbubble
population whose fragility threshold is below said second power
level and with a subsequent third high power signal to measure the
resultant signal.
27. A method according to claim 25 wherein the ultrasound signal
comprises a single pulse train wherein the two differing power
levels are produced such that a first low power pulse is sent at
the beginning of said pulse train and second and third high power
pulses at the middle and the end of said pulse train,
respectively.
28. A method according to claim 25 wherein the ultrasound signal
comprises two or more pulse trains wherein the first pulse train is
of low power and subsequent pulse trains are of higher power.
29. A method according to claim 25 wherein the ultrasound signal
comprises three or more pulse trains wherein the first pulse train
is of low power and each subsequent pulse train is of consecutively
higher power than the previous pulse train.
30. A method according to claim 26 wherein the backscatter
intensity from said microbubbles from the first low power signal is
recorded as a baseline value.
31. A method according to claim 30 wherein the backscatter
intensity from the third high power signal is subtracted from said
baseline to yield a difference value.
32. A method according to claim 31 wherein the difference value is
compared to experimental reference values to measure ambient fluid
pressure.
33. A method according to claim 31 wherein the ambient fluid
pressure measurement is repeated and results are displayed on the
screen of an ultrasound scanner associated with the defined region
of interest.
34. A method according to claim 26 wherein the second high power
signal is at a power level above the fragility threshold of one or
more populations of said composition.
35. A method of determining the fragility thresholds of a plurality
of populations of microbubbles and correlating said thresholds to
ambient pressures at rupture surrounding said microbubbles
comprising the steps of:
a) determining the fragility slope of each said population from
respective curves, said curves determined by measuring acoustic
density as each of said population as it is acoustically
interrogated along a channel versus distance along said
channel;
b) determining fragility curves by plotting each of said fragility
slopes versus mechanical index, a measure of the acoustic power
used to interrogate each said population;
c) identifying the intercept of substantially linear portions of
each of said fragility curves at zero fragility slope as the
mechanical index at the threshold fragility for each said
population;
d) correlating each of said mechanical indices at the threshold
fragility to an ambient pressure from a predetermined mechanical
index-to-pressure relationship.
36. The method according to claim 35 wherein said fragility slopes
are determined for a plurality of populations of microbubbles
having differing wall thickness.
37. The method according to claim 35 wherein said fragility slopes
are determined for a plurality of populations of microbubbles
having differing moduli of elasticity.
38. The method according to claim 35 wherein said fragility slopes
are determined for a plurality of populations of microbubbles
having materials of differing molecular weight.
39. The method according to claim 35 wherein said fragility slopes
are determined for a plurality of populations of microbubbles
having materials of differing hoop strength.
40. The method according to claim 35 wherein said fragility curves
are determined for a plurality of populations of microbubbles
having differing wall thickness.
41. The method according to claim 35 wherein said fragility curves
are determined for a plurality of populations of microbubbles
having differing moduli of elasticity.
42. The method according to claim 35 wherein said fragility curves
are determined for a plurality of populations of microbubbles
having materials of differing molecular weight.
43. The method according to claim 35 wherein said fragility curves
are determined for a plurality of populations of microbubbles
having materials of differing hoop strength.
Description
The present invention relates to a method for measuring real time
ambient fluid pressure within a fluid-filled body cavity using
gas-filled microbubbles and ultrasonic acoustic energy.
BACKGROUND OF INVENTION
Typically, cardiovascular pressures are measured using catheters
which are introduced into the vascular systems via an artery or
vein. Catheters exhibit a finite risk of both morbidity and
mortality with routine usage in the clinical situation. More
recently, sensor tipped guidewires to measure pressure have been
developed. However, this procedure is also invasive with
concomitant associations of morbidity and mortality. There are
currently no adequately accurate direct, noninvasive real time
clinical methods used to measure pressure in the cardiovascular
system. An indirect method exists using Doppler ultrasound as
suggested by Laaban, et al. (Laaban, J., Diebold, B., Zelinski, R.,
Lafay, M., Raffoul, H., and Rochemaure, J., Chest 96, (6):
1258-1262, 1989). With Doppler techniques, blood flow velocities
are measured using ultrasonic scanners operating in the Doppler
mode. By applying the Bernoulli equation and knowing the peak
velocity, it is possible to calculate the pressure drop across a
cardiac valve that created the flow. If one starts measurements in
a vein such as the superior vena cava of known low pressure, one
can calculate the pressure in the right ventricle and pulmonary
artery and even make an estimate of the endiastolic left
ventricular pressure with the technique. The indirect approach is
filled with errors in difficult cases when good data is most needed
and is used only for diagnosing the right side of the heart. Other
non-invasive measurement schemes have been proposed (Blazek;
Vladimir, Schmitt; Hans-J., U.S. Pat. No. 5,447,161; Aakhus, S.,
Soerlie, C., Faanes, A., Hauger, S. O., Bjoemstad, K., Hatle, L.,
and Angelsen, B. A. J., American Journal of Cardiology, 72:
260-267, 1993; Kyriakides, Z. S., Kremastinos, D. T., Rentoukas,
E., Vavelidis, J., Damianou, C., and Toutouzas, P., International
Journal of Cardiology, 33: 267-274, 1991; Neuman, A., Soble, J. S.,
Anagnos, P. C., Kagzi, M., and Parrillo, J. E., Journal of the
American Society of Echocardiography, 11(2): 126-131, 1998) but
show no significant advancement to the field.
In U.S. Pat. No. 3,640,271 to Horton, there is presented the
concept of injecting a single bubble of known size into a patient
for the purpose of measuring blood pressure. The concept was to
stimulate the bubble into resonance ultrasonically and from the
received backscattered signal, determine the resonant frequency of
the bubble. It is further known that if both the diameter of the
bubble and the resonant frequency are known, then the unknown
pressure could be calculated. However, it is not known that the
precision sized bubbles required for the technical approach have
ever been achieved. Also, at present, it would be nearly impossible
to locate bubbles within a specific organ or cavity for the very
low concentration of bubbles required by the technology.
In Pat. No. 4,265,251 to Tickner, the concept of encapsulating a
pressurized bubble within a fused saccharide shell is presented.
The shell begins to dissolve in the circulatory system, thinning
the wall. At some point in its dissolution, the shell fractures and
the bubble escapes and expands. In so doing, it over-expands from
its encapsulated diameter which sets it to free-ringing. A passive
external transducer detects the free ringing signals and, by
applying the same equations identified by Horton, computes the
pressure. A limitation to the technology for usable clinical
practice is the inability to control the point of rupture and the
lack of precision (Osterle S., Sahines, T., Tucker, C., Tickner,
E., et al., The Western Journal of Medicine, 1985 Ott; 143:
463-468.
In U.S. Pat. No. 5,749,364 to Sliwa, the concept for mapping
cardiac pressures is presented by injecting a population of
non-precision microspheres into the blood pool. Theory indicates
that the resonant frequency peak of an encapsulated bubble is
mathematically related to the ambient pressure. By examining the
backscattered signal of the microspheres and from the change in
their frequency spectrum, a map of the pressure in at least two
dimensions is derived. One claimed method for doing this is to
inject two microsphere population types, which exhibit different
backscatter characteristics, and then use these different
characteristics to deduce ambient pressures. Other work in using
frequency shift has been explored. However, no clinical
applications are known to have been developed (Ishihara, K.,
Kitabatake, A., Tanouchi, J., Fujii, K., Uematsu, M., Yoshida, Y.,
Kamada, T., Tamura, T., Chihara, K., and Shirae, K., Jpn. J. Appl.
Phys., 27(Suppl 27-1): 125-127, 1988) possibly due to the
difficulties in measuring in-vivo frequency shifts. Furthermore,
commercial ultrasound scanners have relatively narrow frequency
bandwidths, which would not allow for the frequency range scan
needed to detect changes in resonant frequency, especially in a
formulation with multiple microsphere populations with attendant
multiple resonance peaks.
In PCT No. 98/32378 (De Jong, N., Frinking, P., PCT No. WO
98/32378, Jul. 30, 1998; Bouakaz, A., Frinking, P., De Jong, N.,
Non-Invasive Pressure Measurement in a Fluid Filled Cavity,
Abstract: The Fourth Heart Centre Symposium on Ultrasound Contrast
Imaging, Jan. 21-22, 1999) there is disclosed the use of the decay
of free gas bubbles to measure ambient pressure or temperature. The
decay time of the gas bubble is dependent on the gas type, the
liquid characteristics, the solubility of the gas within the
liquid, the excitation frequency and the ambient temperature and
pressure. However, in their scheme the microsphere is used only as
a transport mechanism which releases the free bubble upon
insonation and the properties of the free bubble are utilized for
pressure measurement. They propose using a series of intermittent
high power pulses to break the capsule of the microsphere,
releasing the gas bubble and then using a series of intermittent
low power pulses to determine the decay time of the bubble and
therefore calculate pressure or temperature. Although the mechanism
described uses a power level to rupture the capsule that is above a
threshold, the fragility or release mechanism of the microsphere
capsule itself is not controlled. Furthermore, if the microsphere
has a very weak capsule, breakage of the capsule can occur at very
low ultrasound powers. This leads to a response which is not
controlled relative to imaging depth and applied power.
SUMMARY OF THE INVENTION
The present invention is directed to a method for measuring real
time pressure in a region of interest in a fluid-filled body
cavity. A composition of gas-containing microbubbles is introduced
into the cavity, the microbubbles having a predetermined fragility
threshold where the fragility threshold is correlated with the
rupture response to fluid pressure, applied acoustic pressure, or a
combination of both fluid and applied acoustic pressures. The
acoustic pressure is applied from an ultrasonic energy producing
source. The composition of microbubbles has predetermined acoustic
response properties correlating to ambient pressure of the
surrounding fluid. When the microbubbles are at the region of
interest, an ultrasonic signal is applied at a power level
sufficient to cause acoustic pressure sufficient to destroy or
disrupt the membrane of the encapsulated microbubble population
having a fragility threshold below the applied power level. Then,
the ultrasound backscatter response is detected from the population
of intact and failing microbubbles remaining at the region of
interest and the backscatter signals are correlated to the
predetermined acoustic response properties of the microbubble
composition to determine the ambient pressure at the region of
interest.
BRIEF DESCRIPTION OF THE DRAWINGS
FIG. 1 is a plot of average acoustic density versus ambient fluid
pressure of microbubbles tested in Example 1;
FIG. 2 is a plot of power Doppler plume length versus ambient fluid
pressure of the test on microbubbles tested in Example 2;
FIG. 3 is a plot of mean acoustic density versus distance, defining
the fragility slope of the microbubbles tested in Example 4;
FIG. 4 is a plot of the fragility slopes of three microbubble
compositions versus mechanical index as described in Example 4 to
define the fragility curve;
FIG. 5 is a graph showing the fragility thresholds of three
different microbubble compositions having three different wall
thickness;
FIG. 6 is a plot of the backscatter signal versus mechanical index
at three different pressures of a microbubble composition having
110 nm wall thickness;
FIG. 7 is a plot of fragility slopes versus mechanical index of a
microbubble composition having 110 nm wall thickness at three
different pressures;
FIG. 8 is a plot of the fragility slope intercept versus pressure
from the data shown in FIG. 7.
DESCRIPTION OF THE PREFERRED EMBODIMENTS
As used herein the term microbubbles is intended to include
microspheres, microcapsules and microparticles which are hollow and
enclosing a core which may be filled with a gas. This may be a
matrix material. It is not necessary for the microbubbles to be
precisely spherical although they generally will be spherical and
described as having average diameters. If the microbubbles are not
spherical, then the diameters are referred to or linked to the
diameter of a corresponding spherical microparticle having the same
mass and enclosing approximately the same volume of interior space
as a non-spherical microbubble. Microbubbles may be comprised of
surface tension stabilized gas bubbles, surfactant stabilized gas
bubbles, lipids and lipisomes, synthetic polymers and biopolymers;
and may further comprise one or more layers of suitable
material.
Ultrasonic backscatter of gas filled microbubbles and bubble
resonance are well known in the ultrasound contrast field. It is
known from the bubble resonance equations that resonant frequency
and pressure are mathematically related. Theory indicates that an
encapsulated gas bubble under insonation has a spectral envelope
with a pronounced peak value at microbubble resonance and this
aspect leads to the prior art methods of using resonant frequency
shifts to determine pressure. However, in simulations with
commercially available microbubble contrast agents, a nearly flat
frequency response has been measured and reported by Boualaz, et
al., (Bouakaz, A., DeJong, N., Cachard, C., and Jouini, K.,
Ultrasonics 36: 703-708, 1998). Thus, using frequency shift, it may
be difficult to estimate in-vivo pressure.
To overcome the microbubble resonance detection problem, this
invention focuses on the encapsulated bubble property of capsule
fragility. The capsule of a microbubble can structurally fail from
both static and dynamic pressure or combinations of the two. The
ultrasonic waves can then disrupt or disintegrate the
freed/escaping gas bubbles sufficiently to decrease their
backscattering cross-section.
The intensity of backscattered signals depends upon the number of
bubbles present in the sample volume and the size of the bubbles.
If the size of the microbubble population is constant, then if
capsules are broken for any reason and the encapsulated gas within
dissolves or disintegrates into smaller bubbles, there is a
concomitant decrease in backscattered signal intensity.
In the case of a gas filled microbubble used in the method of the
present invention, the materials of the capsule wall, and the
properties of those materials are selected and prepared to allow
for the rupture of the capsule at critical or threshold pressures
or within defined pressure limits. The microbubbles prepared in
this manner are referred to herein as having engineered fragility.
By fabricating a microbubble agent with a specifically designed
fragility response and by controlling the applied ultrasound power,
one can selectively rupture a gas containing microbubble, releasing
the gas bubble, which then dissolves or breaks up under insonation
in the ambient fluid. Since backscatter depends upon the presence
of gas bubbles or gas containing microbubbles and their ability to
re-radiate incident signals, backscatter signal intensity is
decreased when the population of microbubbles or free bubbles
decreases or the diameter of the bubbles decreases. However, in
order to alter the backscatter on a clinically useable timescale,
it becomes necessary to both structurally fail the capsule wall and
have the released gas bubble dissolve or disintegrate quickly in
the blood stream. The method herein disclosed takes advantage of
this process by using the decrease in signal intensity and/or
amplitude as a means of identifying the unknown pressure, instead
of the shift in resonant frequency or decay time of a free gas
bubble. Since microbubbles can be seen within cardiac chambers and
even destroyed by ultrasonic imaging scanners, the present
invention provides a viable way for measuring pressure in real time
provided that the microbubbles are designed and fabricated to meet
certain specifications. These specifications concern the
relationship between static and dynamic stresses within the
membrane or capsule shell that when exceeded cause structural
failure. The released free gas bubble should quickly dissolve or
disintegrate into much smaller bubbles so as not to continue to
backscatter incident signals.
Accordingly, having prepared microbubble compositions with
engineered fragility, such compositions are used to measure the
real time ambient fluid pressure at a region of interest in a
fluid-filled cavity, such as the heart. Typically the microbubble
composition will be introduced into the blood system and the
location of the composition within the body can be monitored by
conventional ultrasound scanning. At the region of interest,
acoustic pressure can be applied by a focused ultrasound source to
rupture the capsules of the microbubbles which have a fragility
threshold below the power level of the applied acoustic pressure.
Then the ultrasound backscatter response from the remaining
population of intact microbubbles or disintegrating gas bubbles at
the region of interest is made and correlated to predetermined
acoustic response properties of the microbubble composition to
determine the ambient pressure at the region of interest. The
results may be displayed as a qualitative or quantitative pressure
map using greyscale or color overlays of the base image.
In general, the ambient pressure at a region of interest can be
determined by correlating pressure to the acoustic energy required
to rupture the capsule of the microbubbles. First, a predetermined
correlation of acoustic energy, typically given as a mechanical
index, to pressure is made. This may be experimentally determined
by measuring the microbubble response under pressure at an
acoustically targeted site. Fragility slopes of various populations
may then be experimentally determined from the results of acoustic
density vs. distance along a channel containing an acoustically
interrogated population of flowing microbubbles. The fragility
slopes are then plotted to derive fragility curves of fragility
slope vs. mechanical index. There are substantially linear portions
of each fragility curve which, when extended to the intercept at
the zero fragility slope value, determine the mechanical index at
the threshold fragility. This mechanical index correlates to the
threshold fragility when the microbubbles rupture, and is dependent
upon the ambient fluid pressure.
This can be accomplished by several methods. Since the microbubble
formulations can be made with uniform properties, a fragility
versus pressure characteristic of the composition may be
predetermined by in vitro experiments to develop a response curve.
In one such method when the composition is injected into the
bloodstream, a focused ultrasound scanner is applied at a low power
sound pulse or pulse train to establish a baseline. Then a second
pulse or pulse train is applied at a higher power which is selected
to destroy the segment of the microbubble population whose
fragility threshold is below the power level of the second pulse. A
third pulse or pulse train also at high power is then used to
measure the post destruction backscatter response. The difference
of the backscatter signal from the third pulse versus the baseline
is a measure of the pressure in the ambient fluid and can be
determined by reference to the response curve. The results may be
displayed on the monitor of the ultrasound scanner as a real time
overlay, either in greyscale or color.
Another method of utilizing the microbubbles to measure ambient
pressure is to combine a microbubble composition fabricated from
two or more populations of gas-filled microbubbles where each
population comprises differing fragility thresholds. The fragility
versus pressure characteristics of each population as well as the
combined population are predetermined by in vitro experiments and
the response curve is developed. After administering the
formulation into the bloodstream by bolus injection or infusion, an
ultrasound scanner is focused on the region of interest to apply
successively increasing power pulses or pulse trains. At each power
setting, that population of microbubbles whose fragility threshold
is below the power level of the pulse will be destroyed. The decay
in the intensity of the backscatter signal from each successive
pulse is a measure of the pressure in the ambient fluid. The
pressure measurement is calculated from this decay, referenced to
the known response curve, and is displayed on the monitor.
Yet another method is to inject a single population of microbubbles
having a population with a fragility threshold that is linearly
decreasing with increasing ambient pressure and under constant
applied ultrasound pressure. The fragility curve is determined by
empirical analysis. An ultrasound scan of the cardiac chambers, for
example, may be performed under a power Doppler mode at a constant
power level so that the quantity of microbubble destruction within
the chambers will be dependent upon the ambient pressure. Each
destruction event of a microbubble will result in a decorrelation
signal being detected and displayed by the ultrasound scanner. The
intensity of the Power Doppler ultrasound display will be directly
proportional to the ambient fluid pressure, by way of the
predetermined response.
By yet another method, the pressure in the chamber of the heart or
across a heart valve may be measured. The acoustic density value of
a series of adjacent regions of interest are measured along a
scanline axis of the ultrasound transducer a rapid sequence. The
slope of a plot of the acoustical density versus distance along the
scanline axis is determined and compared to predetermined values of
the acoustical density versus pressure. The data can then be
displayed as a pressure value along the scanline. By repeating the
determination as subsequent scanlines, the pressure value for a
region within the chamber can be determined and displayed.
An alternate method to measure pressure is as follows. A region of
interest containing microbubbles of engineered fragility is
interrogated with a series of ultrasound pulses in rapid succession
of steadily increasing power values: P.sub.1, P.sub.2, . . .
P.sub.n, and respective acoustic density or backscatter signal
levels are determined: B.sub.1, B.sub.2, . . . B.sub.n. A fragility
curve, as shown in Example 4 below, made by plotting the slope
(taken from the results of acoustical density versus distance)
versus mechanical index, is computed to determine the pressure from
either the slope of the fragility curve which occurs in a nearly
linear high power region or from a computed zero slope intercept as
shown in Example 4 below. Comparing the data to a predetermined
table of correlation to pressure will show pressure at the point of
measurement. Real time display of the pressure can then be
accomplished by the ultrasound scanner.
The microbubbles according to the present invention may be a
surface tension or surfactant stabilized gas bubble or have a
mono-layer shell but preferably have at least a bi-layered shell.
The outer layer of the shell will be a biologically compatible
material or biomaterial since it defines the surface which will be
exposed to the blood and tissues within the body. The inner layer
of the shell will be a biodegradable polymer, which may be a
synthetic polymer, which may be tailored to provide the desired
mechanical and acoustic properties to the shell. The cores of the
microbubbles contain gas, typically air, nitrogen or a fluorocarbon
gas. To make the microbubbles rupturable by ultrasound energy, they
must contain a gas to allow acoustic coupling and particle
oscillation. Microbubbles are constructed such that the majority of
those prepared in a composition will have diameters within the
range of about one to ten microns in order to pass through the
capillary system of the body.
Since the microbubbles preferably have an outer and inner layer,
the layers can be tailored to serve different functions. The outer
shell which is exposed to the blood and tissues serves as the
biological interface between the microbubbles and the body. Thus it
will be made of a biocompatible material which is typically
amphiphilic, that is, has both hydrophobic and hydrophilic
characteristics. Blood compatible materials are particularly
preferred. Such preferred materials are biological materials
including proteins such as collagen, gelatin or serum albumins or
globulins, either derived from humans or having a structure similar
to the human protein, glycosoaminoglycans such as hyaluronic acid,
heparin and chondroitin sulphate and combinations or derivatives
thereof. Synthetic biodegradable polymers, such as polyethylene
glycol, polyethylene oxide, polypropylene glycol and combinations
or derivatives may also be used. The outer layer typically has a
chemistry which allows charge and chemical modification. The
versatility of the surface allows for such modifications as
altering the charge of the outer shell, such as by selecting a type
A gelatin having an isoelectric point above physiological pH, or by
using a type B gelatin having an isoelectric point below
physiological pH. The outer surfaces may also be chemically
modified to enhance biocompatibility, such as by PEGylation,
succinylation or amidation, as well chemically binding to the
surface a targeting moiety for binding to selected tissues. The
targeting moieties may be antibodies, cell receptors, lectins,
selecting, integrins or chemical structures or analogues of the
receptor targets of such materials. The mechanical properties of
the outer layer may also be modified, such as by cross linking, to
make the microbubbles suitable for passage to the left
ventricle.
The inner shell will be a biodegradable polymer, which may be a
synthetic polymer. An advantage of the inner shell is that it
provides additional mechanical properties to the microbubble which
are not provided or insufficiently provided by the outer layer or,
enhances mechanical properties not sufficiently provided by the
outer layer, without being constrained by surface property
requirements. For example, a biocompatible outer layer of a
cross-linked proteinaceous hydrogel can be physically supported
using a high moduli synthetic polymer as the inner layer. The
polymer may be selected for its modulus of elasticity and
elongation, which define the desired mechanical properties. Typical
biodegradable polymers include polycaprolactone, polylactic acid,
polylactic-polyglycolic acid co-polymers, co-polymers of lactides
and lactones, such as epsilon-caprolactone, delta-valerolactone,
polyalkylcyanoacrylates, polyamides, polyhydroxybutryrates,
polydioxanones, poly-beta-aminoketones, polyanhydrides,
poly-(ortho)esters, polyamino acids, such as polyglutamic and
polyaspartic acids or esters of polyglutamic and polyaspartic
acids. References on many biodegradable polymers are cited in
Langer, et. al. (1983) Macromol.Chem.Phys.C23, 61-125.
The inner layer permits the modification of the mechanical
properties of the shell of the microbubble which are not provided
by the outer layer alone. For use as an ultrasonic contrast agent,
the inner layer will typically have thickness which is no larger
than is necessary to meet the minimum mechanical requirements, in
order to maximize the interior gas volume of the microbubble. The
greater the gas volume within the microbubble the better the
echogenic properties.
The combined thickness of the outer and inner layers of the
microbubble shell will depend in part on the mechanical properties
required of the microbubble, but typically the total shell
thickness will be in the range of 25 to 750 nm.
The microbubbles may be prepared by an emulsification process to
control the sequential interfacial deposition of shell materials.
Due to the amphiphilicity of the material forming the outer layer,
stable oil/water emulsions may be prepared having an inner phase to
outer phase ratio approaching 3:1, without phase inversion, which
can be dispersable in water to form stable organic phase droplets
without the need for surfactants, viscosity enhancers or high shear
rates.
Two solutions are prepared, the first being an aqueous solution of
the outer biomaterial. The second is a solution of the polymer
which is used to form the inner layer, in a relatively volatile
water-immiscible liquid which is a solvent for the polymer, and a
relatively non-volatile water-immiscible liquid which is a
non-solvent for the polymer. The relatively volatile
water-immiscible solvent is typically a C5-C7 ester, such as
isopropyl acetate. The relatively non-volatile water-immiscible
non-solvent is typically a C6-C20 hydrocarbon such as decane,
undecane, cyclohexane, cyclooctane and the like. In the second
solution containing the polymer for the inner layer, the polymer in
water-immiscible solvents are combined so that the polymer fully
dissolves and the two solvents are miscible with agitation. The
polymer solution (organic phase) is slowly added to the biomaterial
solution (aqueous phase) to form a liquid foam. Typically about
three parts of the organic polymer solution having a concentration
of about 0.5 to 10 percent of the polymer is added to one part of
the aqueous biomaterial solution having a concentration of about 1
to 20 percent of the biomaterial. The relative concentrations of
the solutions and the ratio of organic phase to aqueous phase
utilized in this step essentially determine the size of the final
microbubble and wall thickness. After thorough mixing of the liquid
foam, it is dispersed into water and typically warmed to about
30-35.degree. C. with mild agitation. While not intending to be
bound by a particular theory, it is believed that the biomaterial
in the foam disperses into the warm water to stabilize an emulsion
of the polymer in the organic phase encapsulated within a
biomaterial envelope. To render the biomaterial envelope water
insoluble, a cross linking agent, such as glutaraldehyde, is added
to the mixture to react with the biomaterial envelope and render it
water insoluble, stabilizing the outer shell. Other cross-linking
agents may be used, including the use of carbodiimide
cross-linkers.
Since at this point the inner core contains a solution of a
polymer, a solvent and a non-solvent with different volatilities,
as the more volatile solvent evaporates, or is diluted, the polymer
precipitates in the presence of the less volatile non-solvent. This
process forms a film of precipitate at the interface with the inner
surface of the biomaterial shell, thus forming the inner shell of
the microbubble after the more volatile solvent has been reduced in
concentration either by dilution, evaporation or the like. The core
of the microbubble then contains predominately the organic
non-solvent. The microbubbles may then be isolated by
centrifugation, washed, formulated in a buffer system, if desired,
and dried. Typically, drying by lyophilization removes not only the
non-solvent liquid core but also the residual water to yield
gas-filled hollow microbubbles.
It may be desirable to further modify the surface of the
microbubble, for example, in order to passivate surfaces against
macrophages or the reticuloendothelial system (RES) in the liver.
This may be accomplished, for example by chemically modifying the
surface of the microbubble to be negatively charged since
negatively charged particles appear to better evade recognition by
macrophages and the RES than positively charged particles. Also,
the hydrophilicity of the surface may be changed by attaching
hydrophilic conjugates, such as polyethylene glycol (PEGylation) or
succinic acid (succinylation) to the surface, either alone or in
conjunction with the charge modification.
The biomaterial surface may also be modified to provide targeting
characteristics for the microbubble. The surface may be tagged by
known methods with antibodies or biological receptors.
The microbubbles may also be sized or processed after manufacture.
This is an advantage over lipid-like microbubbles which may not be
subjected to mechanical processing after they are formed due to
their fragility.
The final formulation of the microbubbles after preparation, but
prior to use, is in the form of a lyophilized cake. The later
reconstitution of the microbubbles may be facilitated by
lyophilization with bulking agents which provide a cake having a
high porosity and surface area. The bulking agents may also
increase the drying rate during lyophilization by providing
channels for the water and solvent vapor to be removed. This also
provides a higher surface area which would assist in the later
reconstitution. Typical bulking agents are sugars such as dextrose,
mannitol, sorbitol and sucrose, and polymers such as PEG's and
PVP's.
It is undesirable for the microbubbles to aggregate, either during
formulation or during later reconstitution of the lyophilized
material. Aggregation may be minimized by maintaining a pH of at
least one to two pH units above or below the isoelectric
point(P.sub.i) of the biomaterial forming the outer surface. The
charge on the surface is determined by the pH of the formulation
medium. Thus, for example, if the surface of the biomaterial has a
P.sub.i of 7 and the pH of the formulation medium is below 7, the
microbubble will possess a net positive surface charge.
Alternatively, if the pH of the formulation medium is greater than
7, the microbubble would possess a negative charge. The maximum
potential for aggregation exist when the pH of the formulation
medium approaches the P.sub.i of the biomaterial used in the outer
shell. Therefore by maintaining a pH of the formulation medium at
least one to two units above or below the P.sub.i of the surface,
microbubble aggregation will be minimized. As an alternative, the
microbubbles may be formulated at or near the P.sub.i with the use
of surfactants to stabilize against aggregation. In any event,
buffer systems of the final formulation to be injected into the
subject should be physiologically compatible.
The bulking agents utilized during lyophilization of the
microbubbles may also be used to control the osmolality of the
final formulation for injection. An osmolality other than
physiological osmolality may be desirable during the lyophilization
to minimize aggregation. However, when formulating the microbubbles
for use, the volume of liquid used to reconstitute the microbubbles
must take this into account.
Other additives may be included in order to prevent aggregation or
to facilitate dispersion of the microbubbles upon formulation.
Surfactants may be used in the formulation such as poloxomers
(polyethylene glycol-polypropylene glycol-polyethylene glycol block
co-polymers). Water soluble polymers also may assist in the
dispersion of the microbubbles, such as medium molecular weight
polyethyleneglycols and low to medium molecular weight
polyvinylpyrolidones.
It will be realized that various modifications of the
above-described processes may be provided without departing from
the spirit and scope of the invention. For example, the wall
thickness of both the outer and inner layers may be adjusted by
varying the concentration of the components in the
microbubble-forming solutions. The mechanical properties of the
microbubbles may be controlled, not only by the total wall
thickness and thicknesses of the respective layers, but also by
selection of materials used in each of the layers by their modules
of elasticity and elongation, molecular-weight, hoop strength, and
degree of cross-linking of the layers. Hoop strength being defined
as a mechanical property of a sphere based on the resistance of a
section on the sphere to radial force. Mechanical properties of the
layers may also be modified with plasticizers or other additives.
Adjustment of the strength of the shell may be modified, for
example, by the internal pressure within the microbubbles. Precise
acoustical characteristics of the microbubble may be achieved by
control of the shell mechanical properties, thickness, as well as
size distribution. The microbubbles may be ruptured by ultrasonic
energy to release gases trapped within the capsule into the blood
stream. In particular, by appropriately adjusting the mechanical
properties, the particles may be made to remain stable to threshold
diagnostic imaging power, while being rupturable by an increase in
power and/or by being exposed to its resonant frequency. The
resonant frequency may be made to be within the range of
transmitted frequencies of diagnostic body imaging systems or may
be a harmonic or subharmonic of such frequencies. During the
formulation process the microbubbles may be prepared to contain
various gases, including blood soluble or blood insoluble
gases.
The preferred embodiment is a bi-layered microbubble with a
biopolymer outer shell and a synthetic polymer inner layer.
Typical diagnostic or therapeutic targets for microbubbles of the
invention are the heart, liver, kidney, vascular system and
tumors.
The following examples are provided by way of illustration, and are
not intended to limit the invention in any way.
EXAMPLE 1
A Hewlett Packard SONOS 2500 ultrasonic scanner was used. This
scanner has the capability of measuring the acoustic density (AD)as
a function of time within a region of interest (ROI) displayed on
the video monitor. The scanner was set in the 2D harmonic mode with
send frequency of 1.8 MHZ and receive frequency of 3.6 MHZ. A test
cell was constructed comprising a 3.8 mm diameter cellulose tubing
(the imaging tube)running through a plastic beaker approximately 3
cm below the top. Degassed water was used to fill the beaker. A
flow system was connected to the imaging tube consisting of a
mixing reservoir with microsphere contrast agent suspended in it, a
peristaltic pump and a pressure transducer with digital readout. A
backpressure valve was placed on the drain end of the tube to be
able vary the system pressure. The transducer focus was set on the
center of the imaging tube and the ROI placed within the image of
the tubing lumen. The scanner was set in the AD mode and the AD
readings were recorded by the scanner. The system was run with zero
back pressure to establish a baseline and then the discharge flow
valve was closed to achieve a desired pressure and the procedure
repeated. Pressures were increased to roughly 200 mm Hg and then
decreased during the study.
Several different gas containing microbubble agents were employed
and all yield a linear relationship as exemplified in FIG. 1. Also,
several different power values (mechanical index or "MI")were
utilized. In all cases, a linear decrease of AD was observed as a
function of pressure for a fixed MI, thus demonstrating selective
fragility.
EXAMPLE 2
Hewlett Packard SONOS 5500 ultrasound scanner was used in
conjunction with a an ATS Laboratories, Model 524 Doppler Flow
Phantom. The setup was essentially the same as Example 1. However,
the scanner was set up in the Angio (Power Doppler or Doppler
decorrelation) mode. In this mode, decorrelation events as
determined by Doppler signal processing above a preset threshold
are displayed on the monitor of the ultrasound scanner. The flow
phantom consists of a housing filled with an elastomer which is
designed to mimic the attenuation of living tissue and has four
flow channels of 2, 4, 6, & 8 mm diameter respectively running
through it. The 6 mm diameter channel of the flow phantom was
chosen. The sector scan transducer was oriented along the
centerline of the flow tube. When operated in the Angio mode and
when microbubbles are present in the flow tube of the phantom, a
colored plume derived from the decorrelation events is displayed on
the monitor. The plume corresponds to the disruption or
disintegration of microbubbles within the scanned ultrasound field.
All tests were performed with a pulse repetition frequency of 1.2
kHz, an triggering interval of 2000 ms, 8 pulse Doppler packet and
various values of TIS (Thermal Index, Soft tissue: a measure of
output power) and various formulations microbubbles having an outer
layer of albumin and an inner layer of d, 1 lactide. Using the
caliper function of the system, one can measure the distance from
the first point of insonation to the end of the plume on the
centerline of the channel. Very fragile microbubbles fail instantly
as they enter the sound field and the plume is very short,
typically only a few millimeters, whereas more durable microbubbles
produce a plume which can extend across the image. Since the plume
changes with number of decorrelation events, a measure of, agent
fragility is accomplished and one can then examine the results
under varying ambient pressures. By adjusting the discharge
(backpressure) valve, one can take measurements of the plume length
with varying pressure. If the microbubbles exhibit engineered
fragility properties, the plume length should change under
differing pressure conditions which are dependent upon the
properties of the microbubbles. Indeed this is the case as
exemplified in the FIG. 2.
As can be seen in FIG. 2, there is a linear decrease in plume
length with pressure with a correlation coefficient of 0.98. This
particular test did not require high power levels. The tests were
performed under various power conditions ranging from a TIS value
of 0.2 to 1.0 and in all cases tested there was a linear
relationship between plume length and pressure.
EXAMPLE 3
The set-up as presented in Example 2 above was used with the
exception that the flow phantom was exchanged for a 6.5 cm diameter
by 7.6 cm high acrylic cylindrical phantom chamber with a sealed
top and bottom. A magnetic stirring bar was placed in the chamber
and the chamber filled completely with de-gassed water. Two ports
with luer lock syringe fittings are present to allow connection to
the chamber. One port is connected to the pressure transducer, and
the other port connected through a 3-way valve to a pressurizing
syringe and a sample injection syringe. The S4 transducer of the HP
5500 scanner was placed in a water-filled well on the top of the
chamber, imaging directly downward through a thin acrylic cover.
The HP SONOS 5500 was run in Angio mode and in B-mode Harmonic. The
transducer focus was set at 4 cm in depth, near the mid chamber
position. A diluted sample of gas filled, dual walled microbubbles
(described in Example 8) was injected into the chamber with the
pressure transducer port open to relieve the pressure build-up from
sample introduction. The pressure transducer port was closed and
the 3-way valve switched to the pressurizing syringe, which is
filled with water. The chamber is placed on a magnetic stirrer set
to slow speed.
The ultrasound scanner was turned on and the Power Doppler image
displayed is of a somewhat spherical region of signals from Doppler
decorrelation events within the central region of the phantom. The
pressurizing syringe is then engaged to bring the pressure up to
approximately 150 to 200 mm Hg and back to ambient at approximately
30 cycles per minute. The display shows the region of decorrelation
signals diminishing in size significantly with the increase in
pressure and returning to baseline as the pressure returns to
ambient. The image changes indicate the correspondence of the
Doppler decorrelation signals to the ambient pressure. The results
were recorded to videotape.
EXAMPLE 4
Using the set-up with the HP SONOS 5500 and Doppler Flow Phantom as
detailed in Example 2 above, a test was performed to determine the
variations in fragility thresholds and their relationship to
ambient pressure of microbubbles fabricated with differing wall
thickness. Three samples of microbubbles (described in Example 8)
were fabricated with polylactide inner walls of thickness
approximately 28, 55 and 110 nm respectively. The samples were
interrogated using B-mode Harmonic imaging at 1.8 MHz send and 3.6
MHz receive frequencies. Using the AD function, an ROI was placed
at the input end of the image of the flow channel and densitometry
readings taken at each increment. The ROI was moved laterally, in
increments along the flow direction and readings taken. The
resultant graph, FIG. 3, shows the exemplary results for one sample
(55 nm thickness) as a function of mean AD value versus distance in
centimeters along the flow channel. The slope of this line is
called the Fragility Slope (FS). The results from all samples
indicate a linear decrease in backscattered signal with distance
from the source, and exposure to the ultrasound field.
The test is repeated at MI values from 0.0 to 1.6, in increments of
0.1 MI. The slopes of the resultant measurements are then plotted
versus the power or MI value, as shown in FIG. 4. This curve is
termed the Fragility Curve (FC). All three samples show a region at
low power levels where the slopes are essentially zero, then a
power level where the decrease in signal from the destruction of
microcapsules begins. Determining the curve intercepts with the
zero slope value as shown in FIG. 5, yields fragility threshold
values for the various samples.
Values of AD vs Distance (FIG. 3) were determined at ambient and at
approximately 100 and 200 mm Hg respectively. The y-intercept of
FIG. 3 was taken as an indication of peak backscatter signal. The
peak AD backscatter signal was then plotted versus MI at the three
test pressures in FIG. 6. The curves are roughly equal indicating
that the agent is stable under the test pressures examined and
exhibits the same acoustic behavior at each pressure, and
furthermore that agent concentrations in each test were
comparable.
Fragility Slopes and resultant Fragility Curves for the 110 nm wall
thickness sample where generated at different pressures as above.
Taking the data in the high MI region, from 1.0 to 1.6, wherein the
behavior of the Fragility Curve is nearly linear, three distinct
linear results are seen as in FIG. 7. The intercept of these lines
with the zero slope value is shown in FIG. 8. These curves may be
used to compute the ambient pressure of the fluid.
EXAMPLE 5
Preparation of Gelatin Polycaprolactone Microbubbles
A solution of 1.0 gms gelatin (275 bl, isoelectric point of 4.89)
dissolved in 20 ml deionized water was prepared at approximately 60
C. Native pH of the solution was 5.07. Separately, 1.0 gms
polycaprolactone (M.W. 50,000) and 6.75 ml cyclooctane was
dissolved in 42 ml isopropyl acetate with stirring at approximately
70 C. After cooling to 37 C., the organic mixture was then slowly
incorporated into the gelatin solution maintained at 30 C. and
under moderate shear mixing using a rotary mixer. Once the organic
phase was fully incorporated, the mixing rate was increased to
2,500 rpm for 5 minutes and then stirred at low shear for an
additional 5 minutes. The resulting o-w emulsion was then added
with stirring to 350 ml deionized water maintained at 30 C. and
containing 1.2 ml 25% gluteraldehyde. Immediately after the
addition of the emulsion, the bath pH was adjusted to 4.7. After 30
minutes, the pH was adjusted to 8.3. Low shear mixing was continued
for approximately 21/2 hours until the isopropyl acetate had
completely volatilized. Polyoxamer 188 in the amount of 0.75 gm was
then dissolved into the bath. The resulting microbubbles were
retrieved by centrifugation and washed 2 times in an aqueous
solution of 0.25% polyoxamer 188.
Microscopic inspection of the microbubbles revealed spherical
capsules having a thin-walled polymer shell encapsulating a liquid
organic core. Staining the slide preparation with coomassie blue G
indicated the presence of an outer protein layer uniformly
surrounding the polymer shell.
The particle size spectrum was determined using a Malvern Micro.
Median diameter was 4.78 microns with a spectrum span of 0.94.
EXAMPLE 6
Preparation of Microbubble Agent Formulation
A quantity of microbubbles prepared in a manner similar to example
5 were suspended into an aqueous solution of 25 mM glycine, 0.5%
pluronic f-127, 1.0% sucrose, 3.0% mannitol, and 5.0% PEG-3400. The
suspension was then lyophilized. The resulting dry powder was
reconstituted in deionized water and examined under the microscope
to reveal that the microbubbles now contained a gaseous core.
Staining the preparation with commassie blue G confirmed that the
outer protein layer surrounding the capsules was intact and had
survived the lyophilization process.
Echogenicity was confirmed by insonating at both 21/2 and 5 MHZ a
quantity of lyophilized microbubbles dispersed in 120 ml deionized
water. Measurement was taken at least 15 minutes after dispersion
of the microbubbles to insure that the back scattered signal was
due solely from the gas contained within the microbubbles. The B
mode display showed a high contrast indicating that the
microbubbles were gas filled.
EXAMPLE 7
Preparation of Albumin Polycaprolactone Microbubbles
A 6% aqueous solution was prepared from a 25% solution of USP grade
human serum albumin (Alpha Therapeutic Corp) by dilution with
deionized water. The solution was adjusted to a pH of 3.49 using 1
N HCl. Separately, 8 parts by weight polycaprolactone (M.W. 50,000)
and 45 parts cyclooctane were dissolved in 300 parts isopropyl
acetate at approximately 70.degree. C. Once dissolution was
complete, the organic solution was allowed to cool to 37.degree. C.
With mild stirring, 42.5 gm of the prepared organic solution was
slowly incorporated into 25.0 gm of the albumin solution while the
mixture was maintained at 30.degree. C. The resulting coarse o-w
emulsion was then circulated through a stainless steel sintered
metal filter element having a nominal pore size of 7 microns.
Recirculation of the emulsion was continued for 8 minutes. The
emulsion was then added with stirring to 350 ml deionized water
maintained at 30.degree. C. and containing 1.0 ml of 25%
gluteraldehyde. During the addition, the pH of the bath was
monitored to insure that it remained between 7 and 8. Final pH was
7.1. Low shear mixing was continued for approximately 21/2 hours
until the isopropyl acetate had completely volatilized. Poloxamer
188 in the amount of 0.75 gm was then dissolved into the bath. The
resulting microbubbles were retrieved by centrifugation and washed
2 times in an aqueous solution of 0.25% poloxamer.
Microscopic inspection of the suspension revealed spherical
particles having a thin-walled polymer shell with an outer protein
layer and an organic liquid core. The peak diameter as, determined
by the Malvern Micro particle size analyzer, was 4.12 microns.
The suspension was then lyophilized in a manner similar to that
described in Example 6. The resulting dry cake was reconstituted
with deionized water and examined under the microscope to reveal
that the microbubbles were spherical, discrete, and contained a
gaseous core.
EXAMPLE 8
Preparation of Albumin Polylactide Microbubbles
A 6% aqueous solution was prepared from a 25% solution of USP grade
human albumin by dilution with deionized water. Ion exchange resin
( AG 501-X8, BioRad Laboratories) was then added to the solution at
a ratio of 1.5 gm resin to 1.0 gm dry weight of albumin. After 3
hours the resin was removed by filtration and the pH of the
solution was adjusted from 4.65 to 5.5. Separately, 0.41 gm d-1
lactide (0.69 dL/gm in CHCl.sub.3 : at 30.degree. C.) and 5.63 gm
cyclooctane were dissolved in 37.5 gm isopropyl acetate. The
organic solution was then slowly incorporated into 25.0 gm of the
prepared albumin solution with mild stirring while the mixture was
maintained at 30.degree. C. The resulting coarse o-w emulsion was
then circulated through a stainless steel sintered metal filter
element having a nominal pore size of 7 microns. Recirculation of
the emulsion was continued for 8 minutes. The emulsion was then
added with stirring to 350 ml deionized water maintained at 30 C.
and containing 1.0 ml of 25% gluteraldehyde. During the addition,
the pH of the bath was monitored to insure that it remained between
7 and 8. Final pH was 7.0. Low shear mixing was continued for
approximately 21/2 hours until the isopropyl acetate had completely
volatilized. Polyoxamer 188 in the amount of 0.75 gm was then
dissolved into the bath. The resulting microbubbles were retrieved
by centrifugation and washed 2 times in an aqueous solution of
0.25% polyoxamer.
Microscopic inspection revealed hollow spherical polymer
microbubbles having an outer protein layer and an inner organic
liquid core. The suspension was formulated with a glycine/PEG 3350
excipient solution, then lyophilized. The resulting dry cake was
reconstituted with deionized water and examined under the
microscope to reveal that the microbubbles were spherical,
discrete, and contained a gaseous core.
EXAMPLE 9
PEG Modification of the Microbubble Surface
Microbubbles were prepared in a manner similar to Example 7. After
centrifugation, 4 ml of the microbubbles containing cream
(approximately 11 ml total yield) was resuspended in 31 ml
deionized water. To this was added a 10 ml solution containing 0.3
gm methoxy-peg-NCO 5000 and the pH was adjusted to 8.7. The mixture
was allowed to react at room temperature with mild agitation for
41/2 hours. At the end of this period the pH was measured to be
7.9. The microbubbles were retrieved by centrifugation and washed 2
times in a 0.25% solution of polyoxamer 188. The suspension was
formulated with a glycine/PEG 3350 excipient solution, then
lyophilized. The resulting dry cake was reconstituted with
deionized water and examined under the microscope to reveal that
the microbubbles were spherical, discrete, and contained a gaseous
core.
EXAMPLE 10
Preparation of Wall Modified Albumin Polycaprolactone
Microbubbles
Albumin coated microbubbles were prepared in a manner similar to
Example 7 with the exception that 0.20 gm paraffin was also
dissolved into the organic solution along with the polycaprolactone
and the cyclooctane.
Microscopic inspection of the finished microbubble suspension
revealed spherical particles having a morphology and appearance
virtually identical to those prepared without the addition of
paraffin.
EXAMPLE 11
During a patient examination, the physician injects a bolus of
microbubble pressure agent intravenously. Using the ultrasound
scanner, the physician images the chambers of the heart, primarily
the left ventricle. With the scanner focused within the left
ventricle and electrocardiogram (ECG) leads attached to the
patient, the physician sets the scanner in Power Doppler (Doppler
Decorrelation) mode. The ultrasound scanner is set-up to trigger
based on ECG input, with the trigger point near to or at the end of
diastolic cycle. The intensity of the image is correlation to the
ambient pressure by way of a predetermined response of the
microbubble agent. The resultant image intensity, when compared to
the predetermined response, yields the end diastolic left
ventricular pressure. This information is especially useful in
determine cardiac ejection fraction which is a measure of the
output of the heart.
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